Natural polymers for the microencapsulation of...

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ReviewCite this article: Gasperini L, Mano JF, Reis

RL. 2014 Natural polymers for the

microencapsulation of cells. J. R. Soc. Interface

11: 20140817.

http://dx.doi.org/10.1098/rsif.2014.0817

Received: 23 July 2014

Accepted: 27 August 2014

Subject Areas:biomaterials, biomedical engineering,

biotechnology

Keywords:cell encapsulation, natural polymers,

microencapsulation

Author for correspondence:Joao F. Mano

e-mail: [email protected]

& 2014 The Author(s) Published by the Royal Society. All rights reserved.

Natural polymers for themicroencapsulation of cells

Luca Gasperini1,2, Joao F. Mano1,2 and Rui L. Reis1,2

13B’s, Department of Polymer Engineering, University of Minho, 4806-909 Caldas das Taipas, Portugal2ICVS/3B’s, PT Government Associate Laboratory, Braga/Guimaraes, Portugal

LG, 0000-0003-2133-4379

The encapsulation of living mammalian cells within a semi-permeable hydrogel

matrix is an attractive procedure for many biomedical and biotechnological

applications, such as xenotransplantation, maintenance of stem cell phenotype

and bioprinting of three-dimensional scaffolds for tissue engineering and regen-

erative medicine. In this review, we focus on naturally derived polymers that can

form hydrogels under mild conditions and that are thus capable of entrapping

cells within controlled volumes. Our emphasis will be on polysaccharides

and proteins, including agarose, alginate, carrageenan, chitosan, gellan gum,

hyaluronic acid, collagen, elastin, gelatin, fibrin and silk fibroin. We also discuss

the technologies commonly employed to encapsulate cells in these hydrogels,

with particular attention on microencapsulation.

1. IntroductionCell encapsulation technologies aim at entrapping viable and functional cells

within a semi-permeable matrix. A suitable matrix must be biocompatible, it

must support cell survival and therefore it must be permeable to oxygen, to

the incoming nutrients and to the outgoing toxic metabolites. Other properties

may depend on the specific application. For example, encapsulation techniques

have been investigated for xenotransplantation to treat endocrine diseases, such

as diabetes [1], anaemia [2] and restricted growth [3]. The transplanted cells

must be shielded from the host immune system to avoid the use of immunosup-

pressant drugs that lead to undesirable side effects [4,5]. Therefore, the matrix

should have a fine-tuned porosity to block antibodies and T-cells while permit-

ting the passage of the incoming signalling molecule and the outgoing

response. Furthermore, it should elicit a minimal host reaction and have a

low degradation kinetics to prolong the protection of cells. For applications

other than immunoprotection, the properties of the matrix may be very differ-

ent. For applications directed towards the repair or regeneration of tissue,

biodegradability of the matrix could be desirable. When the matrix degrades,

the entrapped cells may proliferate and create their own extracellular matrix

in place of the artificial one used to entrap them. Given the different appli-

cations of this technology, many biomaterials have been proposed as

entrapping matrix to fulfil the specific requirements. Suitable materials for

cell encapsulation should mimic the extracellular matrix and should be pro-

cessed under conditions compatible with the presence of cells. Many of these

materials are naturally derived polymers that form hydrogels. Hydrogels are

highly hydrated materials composed of hydrophilic polymers that are cross-

linked to form three-dimensional networks. They are soft, highly porous

structures that once implanted induce a low protein adsorption because of

the low interfacial tension with the surrounding fluids [6]. Hydrogels derived

from natural materials have a similar structure to the extracellular matrix of

many human tissues. They are made of polymers similar to the biological macro-

molecules engineered by nature to perform specific functions in a demanding

environment. Some of them are abundant (e.g. marine sources [7]) and further-

more they can often be processed under mild conditions compatible with cell

survival [8].

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perm

eabi

lity

size and shape

degradability

surface properties(protein adsorption/adhesion to tissue,protective shell,bioinstructive motifs,topography)

core properties (internalmicrostructure,hierarchical organization,adhesion motifs)

bulk nature: water content, chemical composition,cross-linking density, stimuli responsiveness,mechanical properties

diffusion of nutrients,oxygen, biomolecules

removal of waste,release ofbiomolecules

cell density,cell clustersor single cellco-culture

Figure 1. A scheme outlining the properties of the microcapsules for the encapsulation of cells.

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Once processed and formed, the hydrogel will exhibit differ-

ent transport properties depending on its structure, chemical

composition and the degree of cross-linking. This aspect is cru-

cial in determining the ability of cells to obtain nutrients and

remove waste that directly influences cell survival or death. In

fact, the hydrogel matrix can act as a mechanical and/or chemi-

cal barrier towards incoming and outgoing molecules. When

the barrier acts as a mechanical obstacle, there is an upper size

limit of the solute that can pass through the matrix. The

corresponding molecular weight cut-off can be quantified by

assessing the permeation of a tracer molecule (e.g. a fluorescent

dye using confocal microscopy) into the matrix or by assessing

outside the matrix the presence of molecules produced by cells.

When the barrier is chemical, the passage of molecules smaller

than the molecular weight cut-off may be hindered or inhibited.

Oxygen, signalling molecules, cell nutrients and metabolites

should be able to diffuse through the matrix without being

blocked. Furthermore, cell survival inside the matrix depends

on permeability for nutrient supply and removal of metabolites,

generating different gradients of concentration along the dis-

tance from the border. Cells that are far from the border of the

matrix will receive a lower amount of nutrients in a given

time. This means that a low mass transfer rate of nutrients

may be enough for cells next to the external border but may

not be for the inner cells. Also, if high cell densities of entrapped

cells are needed, the mass transfer rate of nutrients should also

be higher to sustain them. These observations suggest that

microbeads with high surface area to volume ratio may perform

better in providing the right conditions to support viable

cells [9]. Typically, the upper size limit for these beads is con-

sidered to be 400 mm in diameter, two times the maximum

diffusion distance of oxygen and nutrients from blood vessels

to cells [1]. Microbeads, in fact, guarantee a higher exchange

rate of substances between cells and the surrounding environ-

ment than standard size beads [10]. The smaller size also

contributes to the mechanical properties of the beads—they

are less prone to breakage and, as such, elicit a milder pericap-

sular reaction once implanted [11]. It has been reported that

smaller beads can lead to more inadequately encapsulated

cells that protrude from the beads [12]. However there is also

strong evidence that microsize beads may be more biocom-

patible than standard size beads [13,14]. For these reason

many groups have developed different microencapsulation

techniques that will be the major focus of this review. Some

of the properties of the microcapsules in relation to the

encapsulation of cells are outlined in figure 1.

In the first part of this review, we discuss some selected

naturally derived polymers that can be processed under mild

conditions and that show potential for cell encapsulation

(table 1). The gel phase can be formed by the reversible effect

of external stimuli (temperature, ionic strength) or by perma-

nent cross-linking [40]. We discuss their structure, their

biodegradability [41,42] and their sol–gel mechanism. In the

second part of this review, we describe some of the methods

for cell encapsulation that allow the fabrication of cell-laden

beads and capsules in the micrometric range.

2. Gelling mechanismThe entrapment of cells in a hydrogel usually starts by sus-

pending cells in a water-based solution of a hydrogel

precursor, the sol flowing phase. The suspension then under-

goes a transition to the gel non-flowing phase by physical,

chemical or biochemical processes (generally referred to as

‘hardening’ in this review). Every passage of this process

must be compatible with cell survival and should induce a

minimal stress to cells. This means that, during the entire pro-

cess, the environmental conditions should be as close as

possible to the physiological conditions. Sometimes, however,

it is not possible to comply with these requirements (e.g.

photo-cross-linkable polymers). In these cases, it is important

to note that the toxicity of a particular treatment is related

not only to its nature and intensity, but also to the time of

exposure [43]. In this review, we mainly discuss polymers

that undergo a sol–gel transition in the presence of ions,

when irradiated by light or when the temperature changes.

See figure 2 for a general scheme of these methodologies.


thermalcross-linking

ioniccross-linking

UV

photo-cross-linking

evol

utio

n of

the

gelli

ng p

roce

ss

(b)(a) (c)

Figure 2. A simplified scheme of the three common gelling mechanisms. (a) Thermal gelation: due to a change in temperature the polymer molecules rearrangefrom random coil to helix, then the helices assemble in clusters joined together by the untwined regions. (b) Ionic cross-linking: the sections of the polymerbackbone carrying the charge bind with ions (circles) of opposite charge. (c) Chemical cross-linking (photo-cross-linking): the photoinitiator molecules (greenparticles) in solution form radicals when irradiated by UV and cross-link the polymer chains.

Table 1. Summary of some properties of the materials discussed in this review with references on their use for cell encapsulation. ECM, extracellular matrix.

name source type gelation biodegradability references

agarose seaweed polysaccharide thermal no [15 – 17]

carrageenan seaweed polysaccharide thermal/ionotropic no [18,19]

alginate seaweed polysaccharide ionotropic no [1,20,21]

chitosan crustaceans polysaccharide ionotropic yes [22,23]

gellan gum seaweed polysaccharide thermal/ionotropic not clear [24,25]

hyaluronic acid ECM polysaccharide thermal/photo upon chemical modification yes [26 – 28]

collagen ECM protein thermal/pH variation yes [29,30]

gelatin ECM/collagen protein thermal yes [31,32]

elastin ECM/synthetic protein thermal yes [33 – 35]

fibrin blood protein enzymatic yes [36,37]

silk fibroin silk protein thermal/pH variation dependsa [38,39]aDepends on the processing technique, e.g. b-sheet content.


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Thermoresponsive polymers for cell encapsulation are

polymers that form gels by a change in temperature. Some

polymers form gel when heated above the transition temp-

erature (e.g. elastin and to some extent collagen), while

others form gels when cooled (e.g. gelatin, agarose). Poly-

mers that form gel when cooled present an upper critical

solution temperature (UCST) above which the water and

the polymer are miscible. By decreasing the temperature

below the UCST, the polymer becomes more hydrophobic

and insoluble forming a gel. In contrast, polymers that

form gels upon heating present a lower critical solution

temperature (LCST) below which they are miscible with

water. Most of natural thermoresponsive polymers, in con-

trast with the synthetic ones, exhibit a UCST and form gels by

cooling warm water-based solutions [44,45]. When a polymer

is dissolved in water, three types of interaction are possible

[46]: polymer–polymer, polymer–water and water–water.

The change in temperature makes the polymer–water inter-

action unfavourable, causing, for example, the transition from

random coil to helix to minimize the exposure of the macromol-

ecule to water. The transition of polymers that present an LCST is

driven by a change in the entropy of water (hydrophobic effect),



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while for polymers presenting a UCST the transition is driven by

a change in enthalpy [47]. Thermoresponsive polymers that are

suitable for cell encapsulation should have a sol–gel transition

temperature around physiological temperature.

Some polymers instead form gels in the presence of electri-

cally charged species. These polymers are polyelectrolytes and

carry a net charge along their backbone. When polyelectrolytes

are combined with multi-valent cations of opposite sign they

cross-link forming insoluble complexes. The complexes are

insoluble because the charged groups, responsible for the solu-

bility of the polymer in water, are mutually shielded [48].

Commonly used natural polymers are negatively charged

(alginate, hyaluronic acid, carrageenan), mainly because of

the presence of carboxyl or sulfate groups along the chain.

The influence of the charge type and density on cellular

response is not fully understood. However, many negatively

charged biomaterials do not induce a strong inflammatory

response in contrast to positively charged polymers that tend

to attract inflammatory cells [49].

Ionically and thermally cross-linked hydrogels are phys-

ical gels formed by ionic or secondary forces. These gels can

be dissolved under appropriate conditions such as a reverse

temperature change or when in contact with chelating agents

to remove the cross-linking ions. Other gels are formed by

chemically irreversible cross-linking. Hydrogels formed

by light irradiation are an important family of chemically

cross-linked hydrogels for cell encapsulation. The photo-

cross-linking process involves the presence of photoinitiator

compounds whose chemical nature determines the reaction

rate and the wavelength of absorption [50]. When irradiated

by light, typically in the UV range, the photoinitiators form

free radicals that react with functional groups of the polymer

backbone forming intermolecular bonds. Biopolymers must

be chemically modified to be photo-cross-linkable, typically

by introducing functional groups such as acrylates. Another

family of chemically cross-linked hydrogels that has recently

received a lot of attention is enzymatically cross-linked hydro-

gels [51]. These gels are inspired by the natural cross-linking

reactions occurring in our body such as the formation of

blood clots by the enzyme transglutaminase. The enzymatic

reaction is catalysed at neutral pH and physiological tempera-

ture. Different enzymes can lead to the formation of gels with

different properties [52] and the reaction can be controlled by

controlling the activity of the enzyme [53].

3. Materials3.1. Carbohydrates3.1.1. AgaroseAgarose is a polysaccharide derived from the cell wall of a

group of red algae (Rhodophyceae), including Gelidiumand Gracilaria [54]. The plant is harvested and agarose is

extracted after a series of purification and homogeniza-

tion steps [55]. The main structure of agarose consists of

alternating units of b-D-galactopyranose and 3,6-anhydro-a-

L-galactopyranose. Agarose extracted from different sources

can have different chemical compositions; for example, sul-

fates can be found instead of the hydroxyl groups with a

variable degree of substitution. Agarose is a responsive poly-

mer and its aqueous solutions undergo a sol–gel transition

upon cooling. Above the sol–gel temperature, agarose

exhibits a random-coil conformation in solution, and upon

cooling the structure changes to a double helix. Some of

the helices then aggregate and the hydrogen bonds bet-

ween structural water and galactose stabilize the structure

(figure 2a) [56]. The gelling temperature depends on the con-

centration of the solution, the average molecular weight of

the polymer and its structure. For this reason, there is a

wide range of commercially available agarose, characterized

by different gel strengths and sol–gel transition tempera-

tures. Some of them can be used for cell encapsulation

since their sol–gel transition occurs at around 378C. The ther-

mal sol–gel transition of agarose is reversible and presents a

marked thermal hysteresis, which is a wide temperature

difference between gelling and liquefaction [57].

The average pore size of agarose hydrogels and, as a con-

sequence, the mass transport properties are influenced by the

concentration of the polymer in solution and the settling

temperature. An increase in concentration results in tightly

packed helices that translate to a decrease in pore size [58].

For a Bio-Rad Certified low-melt agarose, Narayanan et al.[59] measured an average pore size of 600 nm for a concen-

tration of 1% w/v decreasing to 100 nm or less when the

concentration was 3%. A decrease in settling temperature

results in gel with smaller pores and higher elastic modulus

(compression test). For example, Aymard et al. [60] showed

a decrease in elastic modulus for a type I-A agarose (Sigma,

368C gelling temperature) from 78 kPa for samples cured at

58C to 53 kPa for samples cured at 358C.

Agarose does not provide adhesion motifs to cells and

does not allow interaction between adherent cells and the

entrapping matrix [61]. However, it can be supplemented

with adhesion molecules of the extracellular matrix, such as

fibronectin [62] or RGD soluble peptide [63].

Agarose is not biodegradable—it can only be degraded

by specific bacteria, not mammals. It can be degraded invitro by agarases, which are classified according to their

cleavage pattern into three types: a-agarase, b-agarase and

b-porphyranase [64–66].

Agarose hydrogels have been extensively investigated for

cartilage repair in vitro [67]. They support chondrocytes in

culture for up to six weeks [68] and gels embedded into the

cells can be placed in a bioreactor that mimics dynamic phys-

iological loading shortly after encapsulation [69]. However, the

poor biodegradability of agarose inhibits the spontaneous

repair process in vivo probably as a result of the foreign

body reaction to this material [70].

3.1.2. CarrageenanCarrageenan is a water-soluble anionic polysaccharide derived

from the Rhodophyceae red algae by alkali extraction. Carra-

geenan is a galactan, like agarose, and it consists of repeat

sequences of b-D-galactose and a-D-galactose with variable

proportions of sulfate groups. In carrageenan, the b-galactose

is D while in agarose it is L. Commercially available carragee-

nan can be divided into three families based on the position

and number of sulfate groups: k-(kappa), i-(iota) and l-

(lambda) carrageenan carrying 1, 2 and 3 sulfate groups,

respectively [71]. Aqueous solutions of k- and i-carrageenan

can reversibly form hydrogels in the presence of cations,

while l-carrageenan does not undergo a sol–gel transition.

In fact, carrageenan in solution has a random-coil confor-

mation and upon cooling k- and i-conformation becomes a



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double helix with the sulfate groups pointing outwards, while

the higher sulfate content of l-carrageenan inhibits the for-

mation of the helicoidal structure [72]. The positive cations

in solution neutralize the charge of the sulfate groups allowing

a tighter aggregation of the helices [73]. Divalent cations are

effective in promoting the formation of strong hydrogels for

k- and i-carrageenan while monovalent ions are particula-

rly effective on k-carrageenan. Hydrogels prepared using

i-carrageenan are softer and more deformable than those

prepared using k-carrageenan [74]. Rochas et al. [75] measu-

red a compression Young’s modulus of 1 kPa for 1.2% w/v

i-carrageenan hardened overnight with a 0.25 M KCl solution,

while Young’s modulus of k-carrageenan was 10 kPa. They

also showed [76] that an increase in Young’s modulus increa-

ses the molarity of the hardening bath and the concentration of

the polymer in solution.

Despite the extensive and documented use of this poly-

mer as an inducer of chronic and acute inflammation

[77–79], there is no consensus on its effect on the inflamma-

tory response of the host when newly developed purified

carrageenan is used [80]. Also, both i- and k-carrageenan

hydrogels have been used for cell encapsulation by ionic

cross-linking [18] and by the formation of complexes with

polycations such as chitosan [81].

Similar to alginate the degradation of carrageenan hydro-

gels is driven by the exchange of ions with the surrounding

fluids, while only some bacteria produce the enzymes that

can cleave the polymer chain [82].

3.1.3. AlginateAlginate is a polysaccharide, a polyanionic linear block co-

polymer containing blocks of (1,4)-linked b-D-mannuroic

(M block) and a-L-guluronic (G block) acids [83]. It is extrac-

ted from many different species of brown seaweed and is also

produced by two kinds of bacteria, Pseudomonas and Azoto-bacter [84]. When derived from seaweed, the crop is cleaned

and then alginate is extracted with a solution of a sodium

salt followed by precipitation [85].

Alginate is a commonly used polymer for encapsulation

of therapeutic agents [86] and ever since the first successful

microencapsulation of pancreatic islets was reported by

Lim & Sun [33], it has become the most studied material

for encapsulation of living cells [1,87]. When multi-valent

cations (e.g. Ca2þ) are added to a water-based alginate sol-

ution, they bind adjacent alginate chains forming ionic

interchain bridges (figure 2b) that cause a fast sol–gel tran-

sition compatible with the survival of the entrapped cells. It

is generally assumed that cations bind preferably to the G

blocks of the chains but relatively recent studies also suggest

that the M block (in particular, the alternating MG block) has

an active role in cross-linking the polymer chains [88]. In algi-

nate, a naturally occurring biomaterial, the relative ratio

between the G and M blocks is not constant and depends

on the seaweed from which it is extracted. The G blocks pro-

vide rigidity to the polymeric structure and the mechanical

properties of alginates are influenced by the ratio of G and

M blocks, and as expected high G alginates result in the for-

mation of stronger gels in compression [89] and tension tests

[90]. Alginates can form polyelectrolyte complexes in the

presence of polycations such as poly-L-lysine or chitosan.

Poly-L-lysine has been widely used to coat the alginate

beads as a way of controlling their molecular weight cut-

off. A positively charged cation may be immunogenic and

attract host inflammatory cells [49,91]. For this reason,

another external alginate coating is often added to the

beads to form the so-called ‘alginate–polylysine–alginate’

(APA) system. However, developments in the characteriz-

ation of APA capsules [92] suggest that these capsules are

not multi-layered; instead they consist of an inner calcium-

alginate core covered by one single external layer of a poly-

L-lysine and alginate blend. The binding strength of the initial

poly-L-lysine layer depends on the relative ratio of the G and

M blocks in the alginate core. Poly-L-lysine does not bind

tightly to alginates with a high content of G blocks because,

in contrast to M blocks, they do not allow complete inter-

action with the polycation. When these capsules are

implanted or incubated they induce a stronger response

than capsules without poly-L-lysine [93,94]. Alginates can

also be combined with other biopolymers to improve the bio-

logical response of the host. Such studies were recently

performed using high-throughput methodologies for the

evaluation of the in vitro [95] and in vivo [96] response to

different combinations of biomaterials. Furthermore, alginate

does not provide cell adhesion motifs, but it can be conju-

gated with RGD peptides to improve cell adhesion [97].

Alginate is characterized by a wide pore size distribution

(5–200 nm) with the most open structure found in alginates

with high G content [98,99]. The permeability of alginate is

strongly influenced by the concentration and nature of the

hardening ions; higher concentrations of ions create tighter

structures (especially in the outer part of the gel in direct con-

tact with the hardening bath) and as a consequence decrease

the diffusion rate of large molecules outside the gel [100,101].

Instead, when the hardening bath consists of salts with low

solubility in water (e.g. CaCo3) the structure that is formed

is more uniform and the hydrogel has higher mechanical

stability [102]. Furthermore, it should be noted that, as most

of the proteins are negatively charged at pH 7, they do not

easily diffuse into the gel while they diffuse out more quickly

than expected [98].

The kinetics of dissolution of the alginate gel in vivodepends on the G/M composition and it is driven by the

exchange of cross-linking ions with monovalent cations of

the surrounding fluids. Once unbounded, the polymer

chain cannot be degraded by the biological activity of the

host. Alginate can be degraded by alginase but this enzyme

is not present in mammals. Alginate degradation, however,

can be manipulated by chemical modifications of the chain.

A slight oxidation of alginate using sodium peroxide leads

to a polymer that degrades in aqueous medium without

interfering significantly in its ability to form gels [103–105].

The degradation kinetics of alginate can also be tuned by con-

trolling its molecular weight distribution by gamma

irradiation [106]. The dissolution of the alginate gel in vitrocan be obtained by the exchange of ions with a buffer (e.g.

phosphate-buffered saline without calcium and magnesium)

or by using chelating agents such as EDTA or sodium citrate.

Being non-adhesive and having been characterized by mild

dissolution methods, alginate hydrogels have recently been

used to manufacture switchable hydrogels, in which alginate

is mixed with cell adhesive matrices such as collagen. In these

constructs, alginate acts as a self-renewal permissive substrate

for human pluripotent stem cells and, when needed, it can be

removed from the matrix permitting the differentiation of

cells on the collagen substrate [107,108].


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Alginate is by far the most studied material for cell encap-

sulation [87], and it has been adopted for many biomedical

applications. Alginate has historically been used as a protec-

tive barrier to enhance cell therapies, for immunoprotection

of pancreatic islets [1,33], treatment of brain tumours [109],

treatment of anaemia [2] and cryopreservation [110].

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3.1.4. ChitosanChitosan is a polysaccharide derived from chitin. Chitin is a

naturally occurring polymer synthesized by many natural

species as a structural component of their exoskeleton [111].

Chitin consists of N-acetyl-b-D-glucosamine chains arranged

in semi-crystalline structures. Chitin is chemically similar to

cellulose, another structural polysaccharide, with a hydroxyl

position at C-2 being replaced by an acetylamide group [112].

Crabs and shrimp shells are the most common sources for

commercially available chitin. To obtain chitin from the

shells, an acid treatment is needed to dissolve the minerals

followed by purification with an alkaline treatment to

remove the proteins retained from the surrounding tissue

[113,114]. The alkaline treatment can also remove some

acetyl groups bound to the amine [115]. When enough

acetyl groups are removed (50% deacetylation or more) the

polymer is called chitosan. In acidic solutions, the primary

amine is protonated and chitosan becomes a soluble, posi-

tively charged polyelectrolyte [116]. As a polycation,

chitosan has extensively been used in a layer-by-layer tech-

nique, which allows the build-up of multi-layered thin films

made of polymers of opposite charge [117,118]. In fact, chit-

osan can easily form polyelectrolyte complexes with other

polyanions such as alginate [119], pectin [120], elastin [121]

or even DNA [122]. For example, chitosan/alginate multi-

layers have been used as an alternative to poly-L-lysine to

coat alginate beads containing cells, producing capsules the

core of which could be liquefied [123,124]. The ability of chit-

osan to form complexes with polyanions is influenced by the

degree of deacetylation because by removing acetyl groups

more free amine groups can participate in the formation of

the complex [125]. The degree of deacetylation influences

the physical, mechanical and biological properties of chito-

san. For example, chitin can be degraded in vivo, by the

activity of lysozymes up to the formation of amino-sugars

[126]. According to Tomihata & Ikada [127], the rate of degra-

dation decreases with the degree of deacetylation, showing a

marked reduction after 70 mol% of deacetylation and being

very low for fully deacetylated chitosan. This is in agreement

with the results of Freier et al. [128], who also reported

prolonged degradation for fully deacetylated chitosan.

Furthermore, they reported a good adhesion and prolifer-

ation for dorsal root ganglion neurons for fully deacetylated

chitosan that decreased proportionally with the degree of de-

acetylation, which was very low at 50%. Also, Chatelet et al.[129] reported that the degree of deacetylation has no signifi-

cant influence on the in vitro cytocompatibility of chitosan

films for keratinocytes and fibroblasts but plays an important

role in the cell adhesion and proliferation of these two cell

lines. Fibroblasts adhere twice as strongly to chitosan as

to keratinocytes. However, while the rate of proliferation of

keratinocytes decreased proportionally with the degree

of deacetylation, the fibroblasts did not proliferate regard-

less of the number of acetyl groups present. The authors

suggested that the high adhesion affinity of fibroblasts for

chitosan materials could alter their growth.

The alkalinization of chitosan acidic solution up to a pH

compatible with the presence of cells (around 7) leads to

the formation of gel precipitates. When the protonated

amine that induces electrostatic repulsion among the chains

is neutralized, a three-dimensional network forms due to

hydrogen bonding and hydrophobic interactions [130]. With

the addition of phosphate salts in solution, chitosan becomes

soluble at physiological pH and becomes thermoresponsive,

forming a gel upon heating to physiological tempera-

ture [131,132]. Chitosan can also be chemically modified

to improve its solubility at neutral pH or its complexation

properties without affecting its biocompatibility and bio-

degradability. These modifications typically involve the

covalent binding of a molecule through graft polymerization

of the free amine groups on deacetylated units or the hydroxyl

groups on the C3 and C6 carbons on acetylated or deacetylated

units [133,134]. A typical example is graft polymerization with

moieties bearing carboxylic groups, such as carboxymethyl.

This modification increases the solubility of chitosan up to

physiological pH without affecting its cationic character [135].

Another example is chitosan grafted with polyethylene glycol

or temperature responsive chains, becoming a water-soluble

thermoresponsive polymer that forms gels at physiological

temperature [136,137].

3.1.5. Gellan gumGellan gum is a polysaccharide derived from the microbial

fermentation product of Sphingomonas elodea. A pure culture

of S. elodea is inoculated in a medium of glucose, nitrogen

and inorganic salts. When the glucose is exhausted by the

biological activity the broth is refined and the gum is col-

lected after precipitation with alcohol [138]. Gellan gum is a

linear anionic polysaccharide composed of tetrasaccharide

repeating units (1,3-b-D-glucose, 1,4-b-D-glucuronic acid,

1,4-bb-D-glucose, 1,4-a-L-rhamnose) [139,140]. Gellan gum is

available in two isoforms, an acylated form that produces

soft hydrogels and a deacylated form that produces hard

and brittle gels. Both forms are in random-coil conformation

at high temperature and upon cooling there is a transition to

double-helix conformation [141]. The double-helices self-

assemble in clusters of anti-parallel structures called junction

zones which are joined together in a three-dimensional

structure by untwined regions of the polymer chain. The

presence of cations in solution is needed to obtain a stable

hydrogel. Without cations the negatively charged carboxyl

side groups repel each other hindering the aggregation of

the helices. Monovalent cations in solution can electrically

shield the carboxyl group by a tighter aggregation of the

helices while divalent cations, in addition to their screening

effect, can bind together two carboxyl groups producing

stronger gels [142,143]. These requirements for the sol–gel

transition (temperature, presence of cations) are compatible

with a physiological environment and as such the sol–gel

transition of gellan gum is compatible with cell entrapment

and survival. Furthermore, the carboxyl group of gellan

gum can be used for chemical modifications to improve its

mechanical properties and stability in vivo. For example,

when modified with methacrylates, the hydrogel obtained

is photo-cross-linkable and presents a higher stability because

of the covalent bonds between the chains [144].



11:20140817

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Ferris & in het Panhuis [145] have characterized the mech-

anical properties of Gelzan, a commercial formulation of

gellan gum. They showed an increase in elastic modulus in

compression from 100 to 400 kPa by increasing the concen-

tration of the polymer in solution from 0.5% to 1.5%. The

number of adherent cells (L-929 mouse fibroblasts) decreased

following an opposite trend and reached a minimum for

gellan gum of 1.5%. They also showed an initial increase in

elastic modulus from 250 to 350 kPa by increasing the con-

centration of calcium ions in the hardening bath from 5 to

10 mM. The elastic modulus then decreased back to 200 kPa

by further increasing the concentration of ions to 15 mM.

The number of adherent cells followed a similar trend and

was maximum for the gellan gum hardened with a solution

of 10 mM.

Some authors have conducted experiments aimed at char-

acterizing the degradation of this hydrogel in vitro and in vivo[146,147]. The degradation rate of the gel is influenced by the

number of acyl groups; in particular, hydrogels made of low

acyl gellan gums degrade more slowly than the high acyl

forms [148]. Furthermore, gellan gum is degraded by the

enzyme galactomannanase [149].

3.1.6. Hyaluronic acidHyaluronic acid is a polysaccharide present in all living

organisms and is found in most connective tissues [150]. It

is a glycosaminoglycan synthesized by membrane-bound

hyaluronan synthases, which distinguishes it from other

glycosaminoglycans that are produced in the Golgi appar-

atus. Hyaluronic acid can be obtained from many tissues by

extraction or enzymatic digestion and it can also be produced

by bacteria. Different sources and extraction protocols can

lead to preparations of hyaluronic acid characterized by

similar molecular weight but different content of endotoxins

and contaminating proteins, leading to a different behaviour

in in vivo and in vitro experiments [151,152]. Hyaluronic acid

is a linear anionic polysaccharide comprising 1,3-b-D-

glucuronic acid and 1,4-b-N-acetyl-D-glucosamine, a structure

conserved in all mammals [153]. It is a hydrophilic polymer

that can form highly viscous solutions at low concentrations

[154]; it is widely used as a lubricant and in preventing post-

surgical adhesions [155]. The molecule of hyaluronic acid in

solution is stiffened by a combination of the chemical struc-

ture of the disaccharide, internal hydrogen bonds and

interactions with the solvent [156]. The polymer has to be

chemically modified to form a hydrogel. This procedure

may involve the modification of the carboxyl or the hydroxyl

group by esterification and cross-linking with, for example,

glutaraldehyde, carbodiimide and divinyl sulfone [157].

For cell encapsulation, the polymer can be treated with

methacrylic anhydride to obtain methacrylated hyaluronic

acid, a photo-cross-linkable polymer (figure 2c) [158]. Seidlits

et al. [159] showed that the mechanical properties of meth-

acrylated hyaluronic acid can be tuned to be close to the

natural central neural tissues. The compressive modulus

depended, as expected, on the initial degree of methacryla-

tion (DM) and ranged from 3 to 10 kPa for a 1% w/v

hyaluronic acid with a molecular weight of 1500 kDa. They

also observed that the diffusion of small molecules (DAPI)

was inversely related to the DM. Bencherif et al. [160]

observed a monotonic increase in the shear modulus as the

DM increased with a variation less pronounced for highly

methacrylated hydrogels. The hydrogel with a DM of 14%

had a shear modulus of 15 kPa, increasing up to 30 kPa for

30% DM and 60 kPa for 90% DM. They measured the influ-

ence of the concentration of the polymer on the shear

modulus for a hyaluronic acid with 32% DM, ranging from

22 kPa for a 2% w/v hyaluronic acid up to 70 kPa for 10%

w/v. They did not observe substantial differences in cell

attachment and proliferation on hydrogels with different DM.

Hyaluronic acid does not favour cell adhesion but it can

be further modified to incorporate adhesion motifs such as

RGD [161]. Hyaluronic acid is used for the encapsulation of

all those cell lines whose extracellular matrix is rich in

glycosaminoglycans or hyaluronic acid itself. For example,

hyaluronic acid has been studied extensively for cartilage

tissue engineering, demonstrating its ability to support chon-

drogenic differentiation of mesenchymal stem cells and the

formation of cartilaginous matrix [162]. Another important

field of application of this polymer is neural tissue engineer-

ing since hyaluronic acid has an important role in the

development of the central neural system, in nerve regener-

ation, in astrocyte activation and proliferation after a spinal

cord injury [163].

Hyaluronic acid is biodegradable in mammals. It is

rapidly degraded by hyaluronidase, b-glucuronidase and

b-N-acetyl-glucosaminidase up to the formation of low mol-

ecular weight hyaluronic acid and oligosaccharides that enter

the glycolytic pathway [150].

Being naturally present in most connective tissue, hyal-

uronic acid is commonly used for the encapsulation of cells

whose extracellular matrix is rich in hyaluronic acid such as

chondrocytes [164].

3.2. Proteins3.2.1. Collagen and gelatinCollagen is the most abundant protein in humans and the

main component of the extracellular matrix of many tissues.

It is mostly synthesized by fibroblasts and osteoblasts [165].

Among the existing collagen types, collagen type I is the

most abundant and is extracted from tissues (ligaments,

skin) by enzymatic and acid treatments [166]. Collagen pro-

teins consist of a unique triple helix extending over a large

portion of the molecule. The helices assemble in complex

supramolecular structures. Every third amino acid of the

chain of the helix is glycine. Glycine, a very small amino

acid, occupies the centre of the helix allowing a tight packing

of the three chains. About one-third of the remaining amino

acids are proline and hydroxyproline and have their side

chains pointing outwards from the helix [167]. Collagen

type I is typically dissolved in diluted acid. Collagen

self-assembles to form a hydrogel when the solution is neu-

tralized (e.g. with NaOH) and heated to physiological

temperature [168]. For cell encapsulation, cells can be mixed

with the neutralized collagen solution and the suspension

can then be moved to an incubator. Collagen contains some

adhesion motifs as RGD (Arg-Gly-Asp) an important tripep-

tide for the interaction between a variety of cells and the

extracellular matrix. Collagen can be rapidly biodegraded in

mammals via collagenases and metalloproteinases to yield

the corresponding amino acids. The rate of degradation can

be controlled by enzymatic treatment or chemical cross-linking

[169]. Because collagen is the main component of the extra-

cellular matrix of many tissues it has found many potential



11:20140817

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applications for tissue engineering [170] such as cartilage repair

[171], mesenchymal stem cell-based therapies [172], bone

regeneration [173], etc. Collagen is the most common clinically

used biomaterial for skin repair [174].

The triple helix of collagen can be broken down into single-

strain molecules to obtain gelatin [175]. Two different types of

gelatin can be obtained by treating collagen: one is the result of

the hydrolysis of the amide groups of asparagine and gluta-

mine into carboxyl groups, while the other is the result of an

acid treatment. The carboxyl groups make gelatin negatively

charged [176]. Gelatin is thermoresponsive, undergoing a

reversible sol–gel transition by cooling a water-based solution

of the polymer below 358C. The hydrogel can be liquefied by

heating it to physiological temperature. This property has been

exploited to fabricate hydrogels with an inner gelatin core that

‘melts’ once placed in physiological conditions [31] or to fabri-

cate porous cell-laden scaffolds with the gelatin beads acting

as porogen [177]. Gelatin can be chemically modified to encapsu-

late cells so that it does not liquefy when placed at physiological

temperature. For example, methacrylate groups can be added

to the side chains of gelatin, resulting in photo-cross-linkable

gelatin–methacrylamide [178].

3.2.2. FibrinFibrin is a major component of blood clots and a key regula-

tor of wound healing. Fibrin is a polymer similar to collagen

formed by the enzymatic polymerization of the protein fibri-

nogen in the presence of thrombin [179]. Fibrinogen is a

water-soluble glycoprotein of 340 kDa, comprising two sets

of three polypeptide chains that are linked together by disul-

fide bonds. Fibrinogen is synthesized primarily in the liver, it

is present in the human blood and its concentration increases

after a trauma. Fibrinogen can be isolated by precipitation

starting from the plasma of autologous blood, for example,

by a series of freezing and thawing cycles or by using chemi-

cals aimed at decreasing the solubility of the protein [180].

Thrombin catalyses the cleavage of fibrinopeptides leading

to fibrin monomers [181]. Cleavage of the fibrinopeptides

occurs in the central N-terminal part of the fibrinogen expos-

ing the ‘A-’ and ‘B-knobs’ binding sites [182]. The knobs

interact with the ‘A-’ and ‘B-holes’ present at the end of the

fibrinogen molecule forming insoluble fibrin fibres. The

branching of the fibres results in a three-dimensional fibrin

network. Finally, in the presence of calcium the transglutami-

nase factor XIIIa cross-links and stabilizes the structure. Factor

XIIIa is derived from the thrombin-mediated cleavage of factor

XIII, a transglutaminase that can promote stem cell adhesion

and proliferation [183]. The formation of the gel and its mech-

anical properties are influenced by the concentration of

fibrinogen and thrombin [184]. Lower concentrations of throm-

bin lead to more compact gels, with thicker fibres and better

mechanical properties. Furthermore, when fibrin is used as

the matrix to encapsulate stem cells, the concentration of throm-

bin and fibrinogen can influence their proliferation rate and their

differentiation. The results from Catelas et al. [185] show that

formulations containing a lower concentration of fibrinogen

(5 mg ml21) can support human mesenchymal stem cell

growth, while a higher concentration (50 mg ml21) has an

increased potential for their differentiation into osteoblasts. In

mammals, fibrin can degrade rapidly owing to the presence of

proteolytic enzymes (fibrinolysis). In fact, during wound heal-

ing, fibrin is gradually degraded and replaced by mature

extracellular matrix. The degradation kinetic of fibrin can be

controlled by protease inhibitors such as aprotinin [186,187].

A cell suspension in fibrinogen and thrombin can be injected

to form a fibrin hydrogel in situ. This particular approach has

found many applications in tissue engineering, specifically for

cell therapies to enhance cell retention after transplant. There

are many examples in the literature of these injectable systems

not only for restoring the myocardium after myocardial infarc-

tion [188–190] but also for other applications such as bone

[191] or muscle regeneration [36].

3.2.3. ElastinElastic fibres are present in all those tissues that require the

recovery of their initial shape after deformation, such as skin,

ligaments and blood vessels [192]. The main components of

elastic fibres are elastin and microfibrils. Elastin is a protein

formed by the polymerization of water-soluble tropoelastin

in a process called elastogenesis. The tropoelastin proteins

are synthesized by cells and they consist of hydrophobic and

cross-linking hydrophilic domains. The hydrophobic domains

mainly consist of valine, glycine and proline, while lysine and

alanine are the main components of the hydrophilic domains

[193]. The tropoelastin proteins are secreted by cells; outside

the wall they self-assemble in a process called coacervation, fol-

lowing a precise patterning of mostly alternating hydrophobic

and hydrophilic sequences [194]. The tropoelastin aggregates are

then deposited and cross-linked onto microfibrillar templates

forming elastin [195].

Natural elastin can be extracted from tissues by harsh

alkaline treatments leading to a poor yield. Typically, elas-

tin-like polypeptides (ELPs) are used, which are synthetic

naturally inspired polymers using human elastin sequences

as their building blocks [196]. ELPs are expressed from a plas-

mid-borne gene in Escherichia coli [197]. The polypeptides

obtained by the purification of the cell lysates are macro-

molecules constituted by large monomers that are repeated

a few times. For this reason and to underline their engineered

nature, Rodrıguez-Cabello et al. [198] proposed the term

elastin-like recombinamer (ELR). The most studied members

of the ELR family are based on the VPGVG pentapeptide, the

most abundant sequence in natural human elastin. Other

ELRs have been synthesized by substituting the fourth

amino acid with other natural amino acids except proline.

The ELRs are thermally responsive and they undergo a

reversible sol–gel transition upon heating. Under the tran-

sition temperature, they are in random-coil conformation

and upon heating the chain folds hydrophobically forming

a regular, ordered b-spiral structure stabilized by hydro-

phobic interactions. The transition temperature is influenced

by the molecular weight, the concentration in solution and

the amino acid composition but can be easily controlled to

be between room and body temperature. Furthermore, biode-

gradation sequences [199,200] or specific cell adhesion motifs

can be added to the polymer chain [201,202]. For example, a

negatively charged ELR containing the RGD motif was pro-

duced by Costa et al. [203] to fabricate microcapsules with a

layer-by-layer technique using chitosan as a polycation.

3.2.4. Silk fibroinSilkworm silk, namely silk produced by Bombyx mori, is a fila-

ment made from two different proteins, sericin (the outer

coating, about 25%) and fibroin. For biomedical applications,


Table 2. A comparison of microencapsulation techniques.

name scalability control on sizeminimum size(d in mm) size dispersion best gelling mechanism

extrusion medium/good medium 80 [212] medium/low fast/ionotropic

lithography medium very good 50 [26] low all

emulsion very good low 10 [213] high/medium thermal

microfluidic medium very good 50 [214] low all

bioprinting medium/low very good 100 [37] low all (more complex with thermal)

superhydrophobic surfaces low very good 1000 [215] low all


11:20140817

9


sericin is removed by a process called degumming, typically

by boiling silk in an aqueous solution of sodium carbonate.

The primary structure of the protein is the recurrent amino

acid sequence (Gly-Ser-Gly-Ala-Gly-Ala)n forming a light

chain (MW 25 kDa) and a heavy chain (MW 350 kDa)

linked by disulfide bonds. The fibroin spun by the silkworm

consists of b-sheet secondary structures arranged so that the

methyl groups and hydrogen groups of opposing sheets

interact [204,205]. This water-insoluble secondary structure,

known as silk II, is stabilized by strong hydrogen bonds

and van der Waals forces. Solvents with high ionic strength

and high concentration of salts (e.g. lithium bromide) are

needed to break down the hydrogen bonds and obtain a

water-soluble random-coil conformation known as silk

I. The solution is then dialysed against distilled water for

3 days to remove the salt. The concentration of the protein

in solution can then be increased by dialysis against poly-

ethylene glycol [206]. Hydrogels can be prepared by fibroin

solutions in water; this process can occur in mild conditions

permitting the encapsulation of cells. The water-soluble silk

I is metastable, and fibroin chains tend to aggregate forming

stable b-sheet structures [207]. The sol–gel transition process

is not fully understood but is known to be influenced by

mechanical stresses, the concentration of the protein, temp-

erature, pH and concentration of salts in solution [208]. For

example, Wang et al. [38] successfully encapsulated cells

into a fibroin gel by sonicating fibroin aqueous solutions at

different concentrations. The gel forming of fibroin can

occur under mild conditions, permitting the encapsulation

of cells. After sonication, the liquid solutions were mixed

with cells and the suspensions were moved to an incubator

where they formed hydrogels. The authors stated that several

physical factors related to sonication, including local temp-

erature increase, mechanical/shear forces and increased

air–liquid interfaces, affect the process of rapid gelation of

silk fibroin (0.5–2 h). Kim et al. [207] characterized the influ-

ence of temperature, Ca2þ and Kþ concentrations, and pH on

the sol–gel transition of fibroin. They showed a decrease in

gelation time at 378C from 30 days to one week by increasing

the concentration of the polymer in solution from 2% to 20%

and that the presence of ions could shorten the time needed

for the sol–gel transition even more. They also measured

the compressive modulus of the gels at 378C, which increased

from a few kPa for fibroin 4% up to 4 MPa for fibroin 16%.

They also observed a decreased average pore size for the

gels made from higher concentration of polymer. Fibroin

hydrogels can also be prepared with low cytotoxic chemi-

cal cross-linkers such as genipin. Sun & Incitti [209]

characterized genipin cross-linked fibroin hydrogels formed

after 48 h of incubation at 378C. The long incubation time

was necessary to obtain a high cross-linking degree (about

93%) needed for the formation of the gel, since genipin pre-

ferentially reacts with lysine and arginine which are present

in fibrin only in low percentages. The hydrogel obtained

had a storage modulus of 41 kPa and supported the long-

term proliferation of pluripotent cells.

Fibroin undergoes a slow proteolytic degradation and the

protein is broken down into smaller polypeptides and free

amino acids [210]. The rate of degradation depends on the

b-sheet content and is related to the preparation methods of

the hydrogel [211].

4. MethodsThe materials and the hardening strategies should be combined

with adequate technologies to encapsulate cells. This section will

explore methods that allow the production of hydrogels with

controlled size, often in the form of capsules or particles in

the micrometre range (table 2). Most of these capsules or parti-

cles are spherical since there is a wide range of applicability

of spherical-based devices for tissue engineering and other

biomedical applications [216].

4.1. ExtrusionThe easiest method to encapsulate cells into hydrogel beads is by

gravitational dripping. A suspension of hydrogel precursor and

cells is extruded through a small tube (i.e. a needle), the drop

grows and when it reaches a critical mass it freely falls into a suit-

able hardening bath. According to Poncelet et al. [217], the final

diameter of the capsule is influenced by the density of the sol-

ution, the surface tension of the drop and the diameter of the

pendant droplet neck (approximately the external diameter of

the needle). This method usually leads to capsules above 1 mm

diameter and because of the impact onto the bath they are not

always spherical. Pereira et al. [24] applied this method to sus-

pensions of immortalized mouse lung fibroblasts and gellan

gum with two different degrees of acylation, using a phosphate

buffer as the hardening agent. The beads had a tear shape and

there was a significant difference in diameter between the

beads obtained using the two polymers due to the different

properties of the solutions. They did not observe any substantial

difference in cytotoxicity and cell viability up to 72 h.

When the extrusion is made directly into a hardening bath

without any air gap, the interfacial shear forces on the drop at

the tip of the needle inhibit its growth. This technique produces

fibres and is called wet spinning. Popa et al. [218] compared the

two techniques showing that beads fabricated by dripping were


V

0 ms 0 ms

(b)(a)

Figure 3. (a) Scheme of the bio-electrospray encapsulation. A potential is applied to the metallic needle of a syringe placed above a metallic plate connected to theground. When the potential applied is zero the drops freely fall from the needle as in gravitational dripping. By increasing the potential, there is a transition to ajetting mechanism and small mono-sized beads are obtained. (b) Real images of dripping (left) and jetting (right) modes. Adapted from Xie & Wang [224] withpermission (Copyright & 2007, Elsevier).


11:20140817

10


at least two times bigger than wet spun fibres. The hydrogels

were used to encapsulate ATD5 cells into alginate–carrageenan

hydrogels. This mixture of hydrocolloids is interesting since the

two polymers share a similar gelling mechanism. In fact, they

both cross-link in the presence of positively charged cations

and a single bath of CaCl22KCl can be an adequate hardening

agent for the mixture.

Smaller beads can be obtained by applying different methods

aimed at ‘breaking’ the jet. One of these methods is by coaxial air

flow, where a compressed gas is forced around the extruding

droplet and the shear stresses applied by the gas cause the dro-

plet to detach before it would fall due to gravity. Higher gas

flows produce smaller beads while higher viscosities of the sol-

ution have a stabilizing effect resulting in bigger beads [219].

Using this encapsulator, Zhou & Xu [220] successfully encapsu-

lated human umbilical cord mesenchymal stem cells in a blend

of oxidized alginate and fibrin. The microbeads (several hundred

micrometres) showed a quicker degradation rate and a higher

cell density than pure and oxidized alginate. This device has

also been used by Mazzitelli et al. [221], who adopted this tech-

nique to encapsulate IB3-1 cells in alginate using a bath of

barium chloride as the hardening agent. The microencapsulation

procedure did not alter the viability of the encapsulated cells and

the cells could still be induced to pro-inflammatory responses

after treatment with tumour necrosis factor-a. A coaxial encapsu-

lator can also be used with a liquid instead of a gas to break the

drop. Sakai et al. [222] used such a device to entrap cells in a

multi-layered capsule made of phenolic-derived alginate and cel-

lulose. The extruding liquid was a suspension of the polymers,

horseradish peroxidase and cells. In this case, the outer fluid

was liquid paraffin that contained hydrogen peroxide to allow

the enzymatic cross-linking of the cell suspension.

There are other ways to break the drop and create microcap-

sules, one of which is the vibrational encapsulator, characterized

by Mazzitelli et al. [212]. This system is based on the principle

that laminar liquid jets break up into equally sized droplets by

applying a vibration to the extrusion nozzle. By using low flow

rates and high frequency of vibration, it is possible to obtain

beads up to 100 mm. They were able to produce calcium-alginate

gel beads with a very narrow size distribution and excellent

morphological characteristics. The microcapsules did not alter

the morphology, viability and functional properties of neonatal

porcine islets.

Prusse et al. [223] have extensively characterized the JetCutter,

a droplet generator that can be used with highly viscous fluids.

This technology is based on a mechanical cut of a liquid jet by

rotating cutting wires. The droplet size mainly depends on the

diameter of the nozzle, the rotation and frequency of the cutting

tool, the number and the diameter of the wires. The diameter of

the beads produced by this technology ranges from 200 mm up

to a few millimetres. With the JetCutter, Schwinger et al. [20]

successfully encapsulated murine fibroblasts into alginate/poly-

L-lysine complexes, showing that encapsulated cells were able to

survive in culture for extended periods of time with unchanged

rates of proliferation and preserved morphology.

Smaller beads can also be obtained by applying an electrical

potential to a metallic extruder (figure 3). The technique, com-

monly referred to as bio-electrospray, consists in spraying a

polymer solution pumped through a needle connected to a

high-voltage generator [225]. The solution at the tip of the

needle reacts to the presence of the electrical charge by accumulat-

ing charges of opposite sign on its surface. The resulting

Coulombic repulsive forces create stresses on the liquid surface

and deform it into a conical shape known as the Taylor cone

[226]. If the electric field is sufficiently high, this electrostatic stress

overcomes the surface tension at the apex of the liquid cone and a

jet of drops is formed to expel the excess of surface charge (Rayleigh

limit) [227]. Once the jetting condition is initiated the size of the drop

is independent of the voltage applied and the distance from the

target anode. Gasperini et al. [21] used this technique starting

from a suspension of B50 mouse cells and alginate. The cell-laden

droplets of the jet were collected in a calcium chloride bath acting

as the hardening agent. They showed that, by increasing the concen-

tration of cells in the suspension, higher applied voltages were

needed to initiate the jetting condition. However, once the condition

was initiated there were no substantial differences in the diameter of

the beads, which ranged from 200 mm up to few millimetres. Cells

did remain viable up to 30 days after encapsulation but they

showed a different behaviour from control cells once released

from the beads and cultured on a tissue culture plate. Cells migrated

to form clusters leaving large portions of completely cell-free tissue

culture plate and were not able to reach confluence.

4.2. LithographyLithography techniques for cell encapsulation can be divided

into soft lithography and photolithography [228] (figure 4).

Soft lithography is a strategy based on self-assembly and replica

moulding for carrying out micro- and nanofabrication [229]. In

this technique, a polymer solution is poured or spin-coated

onto a master and then is cross-linked, obtaining a rubbery

replica. The replica can contain channels that can be filled with


(b)

(a)

Figure 4. (a) Soft lithography approach to cell encapsulation. A mould is placed onto a cell suspension and the microvolumes are filled by capillaryforces. The mould, transparent to UV, is moved onto a clean surface and placed under UV. The microgels that form can be detached from the mould. (b) Photo-lithography. A mask is placed on the cell suspension and then is irradiated by light for cross-linking. Microgels are formed and the rest of the suspension can bewashed away.


11:20140817

11


a suspension of a hydrogel precursor and cells. The channels are

usually filled by pressing the replica onto the cell suspension. This

technique was adopted by Khademhosseini et al. [26] to encapsulate

NIH-3T3 in photo-cross-linkable hyaluronic acid using a polydi-

methylsiloxane (PDMS) replica. They filled the voids in the

replica by pressing it onto the cell suspension, then the replica

was moved onto an acrylated glass substrate and exposed to UV.

After 6 h, about 85% of the cells were viable and after 5 days cells

were able to emerge from the hydrogel.

In photolithography, a mask is placed on a suspension made

of cells and a photo-cross-linkable hydrogel precursor. The sus-

pension is then irradiated by light, typically in the UV range.

The area of the suspension accessible by light cross-links forming

a hydrogel while the uncross-linked portion of the suspension can

be washed away. The hydrogels obtained in this way are thin and

their heights can be increased by repeating the process, thus

obtaining multi-layer structures [230]. Using this technique, Shin

et al. [231] encapsulated NIH-3T3 cells into hybrid hydrogels of

photo-cross-linkable gelatin and graphene oxide on glass slides

coated with polyethylene glycol diacrylate. They optimized the

photo-cross-linking process by increasing the time of exposure

when a higher concentration of graphene oxide was present in

the gel. The hybrid hydrogels demonstrated tunable mechanical

strength and enhanced electrical properties. The gels supported

cellular spreading with improved viability and proliferation.

4.3. EmulsionThe preparation of spherical objects by emulsion-based techniques

has been widely used in the pharmaceutical industry [232]. Encap-

sulation by emulsion is usually obtained by dispersing a hydrogel

precursor into a non-miscible phase. Surfactants can be used

to stabilize the emulsion and to obtain smaller drops (micro-

emulsion) [233]. When the dispersion reaches equilibrium, the

polymer drops are hardened according to the sol–gel mechanism

of the hydrogel. Luan & Iwata [16] adopted this technique to

encapsulate rat islets in agarose microbeads carrying complement

receptor-1, a potent inhibitor of complement activation pathways.

A suspension of islets and agarose was mixed with liquid paraffin

at 408C. The tube containing the emulsion was then immersed in

an ice bath for 3 min to harden the agarose beads. After adding

buffer to the tube and stirring at 48C, the beads could be collected

and the complement receptor was immobilized on their surface. It

was shown that agarose microbeads functionalized with the

complement receptor contributed to the marked prolongation of

islet graft survival in rat-to-mouse xenotransplantation compared

with plain agarose microbeads. Batorsky et al. [234] used a col-

lagen–agarose hydrogel to encapsulate adult mesenchymal stem

cells using PDMS as a non-miscible phase. This mixture of poly-

mers is challenging since both polymers present a sol–gel

transition at physiological temperature but with an opposite

behaviour: one is a hydrogel when the other is a fluid. Collagen

was neutralized at 48C with NaOH at 378C followed by mixing

it with agarose at 608C. The mixture was first added into the emul-

sification bath at 378C, stirred for 6 min and then it was cooled

with an ice bath for 30 min. They obtained beads from 30 to

150 mm in diameter, containing viable cells up to 8 days. They

also showed that by increasing the relative concentration of col-

lagen in the mixture, as expected, cells were more spread due to

the affinity for collagen.

4.4. MicrofluidicsMicrofluidics is a technique dealing with the handling of fluids in

microenvironments, such as microchannels where the flow of

fluids is generally laminar. The flow is characterized by low Rey-

nolds numbers [235], meaning that it is dominated by viscous

stresses with negligible inertia effects. The laminar flow allows

a fine control over the characteristics of the microdrop [236].

The generation of the microdrops usually involves the formation

of emulsions of the polymer droplets in a non-miscible con-

tinuous phase. The generation of individual drops through

microfluidics can be seen as a bottom-up approach to emulsi-

fication, compared with standard emulsification techniques

being the top-down approach. In this bottom-up approach, a

suspension of hydrogel precursor and cells is injected into

a microchannel and the droplets are formed when the suspen-

sion intersects the continuous phase coming from other inlets.

The intersection between the channels can have different geo-

metries, with the T-junction [214,237] and the flow-focusing

[238,239] being the most common (figure 5).

When two channels merge at a right angle the junction is

called a T-junction. In this set-up, the continuous phase flows

in the bigger main channel while the cell suspension is injected

into the orthogonal inlet. The suspension coming from the

orthogonal inlet penetrates the continuous phase and the droplet

begins to grow. The flow intensity and the pressure gradient in

the main channel bend the drop towards the direction of the


(b)(a)

Figure 5. Microfluidic encapsulation devices, real images and schemes showing the hydrogel precursor and the continuous phase. (a) T-junction, hydrogel precursorfrom the bottom channel (adapted from Fu et al. [240] with permission. Copyright & 2010, Elsevier). (b) The flow-focus, hydrogel precursor form the left channel(adapted from Anna et al. [239] with permission. Copyright & 2003, AIP Publishing LLC).


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12


flow until it breaks. Its size is influenced by the size of the micro-

channel and the flow intensity of the two phases. In the flow-

focusing set-up, the continuous phase flows in side channels

while the cell suspension flows in an inner channel between

them. At the junction, the continuous phases force the cell sus-

pension through an orifice located downstream of the inner

channel. The inner fluid then breaks forming a droplet. The

beads can be hardened (e.g. by UV or thermal gelation) inside

or outside the microfluidic chip. Sakai et al. [31] proposed a

flow-focusing microfluidic device able to fabricate cell-enclosing

gelatin capsules in a two-step process. First, they fabricated

microparticles of less than 200 mm and then used the same

device to coat the particles with gelatin incorporating phenolic

hydroxyl groups. After cross-linking the gelatin coating via a

horseradish peroxidase catalysed reaction, a hollow-core struc-

ture was obtained by liquefying the inner unmodified gelatin

after incubation at 378C. Jun et al. [241] reported an encapsulation

technique that produces microfibres and is based on a flow-

focusing device using a cross-linking agent as the continuous

phase. In this experiment, they used a suspension of pancreatic

islets in an alginate–collagen solution for the inner channel

and delivered a calcium chloride solution through the outer

channels. In this way, they obtained a coaxial flux of cell suspen-

sion and hardening agent inside the cylindrical outlet. The

extruded fibres were transplanted and the entrapped islets sur-

vived for more than four weeks. The fibres did not degrade

and successfully protected the cells from the host immune

system. They hypothesized a synergistic effect between alginate

hydrogel (acting as an immunological barrier) and collagen

(supporting the viability of the cells).

4.5. BioprintingBioprinting is a rapid prototyping technology that consists in the

computer-aided layer-by-layer deposition of cells [17,242].

Recently, many bioprinting techniques have been developed,

the working principle of which is often similar to standard

printers for documents. In conventional printers, the ink is

transferred from a reservoir or a substrate and deposited onto

the paper by a piezoelectric/thermal head or a laser beam. The

paper is engineered to absorb the ink. Similarly in some

bioprinters, a suspension of cells, called bio-ink, is deposited

by the printer onto a specific substrate, called bio-paper. The

bio-paper is often engineered specifically for the bio-ink (to

cross-link the bio-ink or to be cross-linked by it). Some bioprin-

ters can be seen as an encapsulating device on top of a

positioning system; in fact, bioprinting is often referred to as a

bottom-up approach using microdrops as building blocks [243].

Cui & Boland [37] adopted this technique to print cell-laden

fibrin scaffolds. In this case, they used bio-ink of thrombin and

human microvascular endothelial cells that was printed using

a modified commercial printer for documents. The cells could

proliferate and connect to each other to form a confluent lining

along the fibrin scaffold after 21 days of culture. Bioprinting

technologies rely on a high-resolution placing system, and this

allows unusual encapsulation methods. For example, it is poss-

ible to deposit a hydrogel precursor and then dispense a cell

suspension on the exact same spot. Lee et al. [29] successfully

encapsulated cells in collagen by depositing collagen type I on

top of a substrate coated with sodium carbonate, acting as a

pH-altering cross-linking agent. Then the droplets of cells in cul-

ture medium were dispensed on the partially cross-linked

collagen layer to be lodged inside. Finally, the hydrogel was

cross-linked by a ultrafine mist of calcium carbonate solution.

More recently, Gasperini et al. [244] used a bio-electrospray

deposition system to print cell-laden alginate scaffolds on top

of a gelatin-enriched biopaper. Using this hydrogel as biopaper,

it was possible to cross-link the alginate upon contact onto the

substrate and print samples up to 3 mm thick. Billiet et al.[245] used a commercial bioprinter to fabricate cell-laden con-

structs using a photo-cross-linkable gelatin. They adapted the

printer to obtain a homogeneous temperature during the extru-

sion of the liquid gelatin and with a Peltier cell they cooled the

deposition substrate below the sol–gel temperature. Further-

more, by optimizing the printing and UV hardening processes

they obtained highly ordered three-dimensional scaffolds with

a high viability of entrapped cells.

4.6. Superhydrophobic surfacesA drop of water placed onto a superhydrophobic surface will

maintain a spherical shape. Cells can be encapsulated by drip-

ping onto a superhydrophobic surface a suspension of cells in

a water-based hydrogel precursor [215]. Then the drops, almost

completely surrounded by air (or another desired atmosphere),

are hardened over the hydrophobic surface maintaining the

spherical geometry. This technique permits the preparation of a

large range of particulate systems under mild conditions avoid-

ing any loss of the cargo during the processing (encapsulation

efficiency very close to 100%). Moreover, there is a good control

over the size of the particles dictated by the volume of the dis-

pensed suspension. Lima et al. [22] adopted this technique



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13


using a suspension of thermoresponsive chitosan (solution of

chitosan and glycerophosphate) and cells. The drops were

placed on a superhydrophobic surface of polystyrene and a

hydrogel complex was formed by the addition of sodium tri-

polyphosphate. Upon heating, at 378C a second sol–gel

transition could be observed, predicting a hardening effect

when exposed to physiological conditions. Cells could remain

viable up to 7 days. Using this method, mesenchymal stem

cells were encapsulated in alginate beads containing fibronectin.

Cells exhibited good viability and osteogenic capability. More-

over, in vivo tests showed that these particles could stimulate

bone regeneration in cranial defects [246]. More sophisticated

layered particles could be obtained by the sequential deposition

of hydrogel precursors over previously hardened particles allow-

ing the distribution of cells or drugs in compartments with a

radial arrangement [247]. Future developments on this technol-

ogy should include dispensing systems able to place smaller

volumes of hydrogel precursors over adequate superhydropho-

bic surfaces to obtain particles smaller than 800 mm.

4.7. CommentsExtrusion techniques range from simple semi-automatic extru-

sions of hydrogel precursors into a hardening bath to more

complex patented systems. Generally, these techniques are best

suited for fast gelling hydrogels, such as alginate. With these

polymers, the surface of the droplet quickly cross-links once in

contact with the hardening agent and the hydrogel coating

that is formed stabilizes the shape of the drop. The liquid core

can then be hardened by keeping the capsule in the bath, allow-

ing the hardening agent to diffuse inside the capsule. The most

interesting extrusion techniques are those able to fabricate gels

in the micrometre range. Although these techniques are able to

produce beads with a relatively good size distribution, it is

not trivial to predict the final size of the beads. If we take

as an example the bio-electrospray process and alginate, the

beads are formed from a suspension of a shear thinning non-

Newtonian fluid and viscous particles (cells) under the effect

of an electric field. The final size of the droplets depends on

the diameter of the needle, the electrical and rheological proper-

ties of the suspension, the flux of the suspension and the applied

voltage. General qualitative information can be drawn but it is

difficult to predict quantitatively the influence of these par-

ameters on the final outcome. A similar observation can be

made for other extrusion techniques. Techniques other than

extrusion, such as lithography, have a better control on size,

meaning that the final dimension of the bead can be easily pre-

dicted. In fact, in soft lithography the droplet is physically

constrained inside a mould of predefined dimensions while in

photolithography the dimensions of the microgel are determined

by the dimensions of a photoresistant mask. Lithography tech-

niques, specifically soft lithography, rely on the production of

masters or masks that involves technologies which are not avail-

able to all research facilities, while extrusion or emulsion

techniques rely on much more common equipment. The same

can be said for microfluidic techniques since the microfluidic

chip is usually fabricated by lithography or advanced microma-

chining. In microfluidics, there is also an added layer of

complexity represented by the flow of the fluid in the microchan-

nels. The chip has to be properly designed to accommodate

inlets, outlets and the junctions where the droplet is formed by

the intersection of the fluids. Furthermore, pumps are needed

to control the flux of fluids inside the channels. However, once

the chip set-up is optimized, microfluidic techniques allow an

easier continuous encapsulation of cells with minimal human

intervention. While photolithography techniques are restricted

to photo-cross-linkable polymers, lithography and microfluidics

are really versatile and allow the encapsulation of cells with basi-

cally all the polymers (and their modifications) discussed in this

review. In soft lithography, once the fluid has filled the channel,

it is trivial to put it in contact with hardening agents, change its

temperature or irradiate it by UV. Similarly, in microfluidics

an inlet can carry the cross-linking agent, the chip can be

heated or cooled and it can be manufactured with mate-

rials transparent to UV such as PDMS, a common material to

manufacture microfluidic chips.

Bioprinting might be the more complex technique to encap-

sulate cells. It is based on a deposition/encapsulation system

which is coupled to an accurate positioning system. Even if com-

plex, bioprinting can rely on cheap but accurate machines

developed for the computer market (document printers and

their controlling drivers). These printers make this technology

appealing in the face of little investment. Furthermore, they are

colour printers and usually contain four reservoirs (red, green,

blue and black) to allow, for example, the encapsulation of differ-

ent kinds of cells suspended into different hydrogel precursors.

These systems are particularly complex encapsulating machines

but this complexity is justified by the possibility to design

drop-by-drop arbitrary structures that translate directly into

cell-laden scaffolds. On the opposite side are emulsion tech-

niques, which are simple, based on common equipment and

can be easily scaled to produce a high number of cell-laden

beads. These techniques can also rely on a well-established

industrial technology developed for cosmetics and pharmaceuti-

cals. The droplets obtained by emulsion typically have a wider

size distribution than those obtained by other techniques. How-

ever, water-in-oil emulsion can be stabilized using surfactants

(i.e. Tween and Span) that may help in narrowing the distri-

bution and controlling the size of the droplets. Also in this

case, as with extrusion, predicting the final dimension of the

beads is not trivial. The system has to be carefully designed by

choosing the surfactants and their relative amount in the suspen-

sion, considering the specific hydrogel precursor and the apolar

phase used. The relative concentration of each component is

usually established empirically. When thermal treatment is

used to harden the droplets careful attention should be given

to the kinetics of the sol–gel transition. Typically during this

treatment, the suspension of droplets and non-miscible phase is

quickly cooled (e.g. with an ice bath) to induce the sol–gel

transition. The thermal treatment should be long enough to

obtain the hydrogel but not too long to cause thermal shock

to cells.

5. ConclusionSince the first cell encapsulation by Lim & Sun [33], a notable

amount of research has been aimed at optimizing the

materials and process selection. Many authors have propo-

sed materials and methods tailored for specific applications

that range from the encapsulation for xenotransplantation

to the production of scaffolds. Among these materials, an

important part is represented by naturally derived polymers

owing to their inherent similarity to the extracellular matrix

and mild processing conditions. In this review, we descri-

bed how some of these authors have combined the

knowledge of specific polymers and their sol–gel mechanism

with specific processing conditions compatible with the pres-

ence of cells. These experiments could be a model to further

develop and improve this technology that could have a huge

impact in the field of biotechnology and other biomedical

applications.

Funding statement. This work was supported by European ResearchCouncil grant agreement ERC-2012-ADG 20120216-321266 for projectComplexiTE.


14


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